Computed tomography (CT) is a method by which it is possible to produce a 3-dimensional reconstruction of the distribution of tissue densities or, more precisely defined, a 3-dimensional distribution of the linear attenuation coefficients for X-ray irradiation of the interior of the body, with a high spatial resolution and a high quantitative precision in the resolution of the densities.
It is technically possible nowadays to map and reconstruct a large volume region with a short exposure time. To do so, use is made mainly of flat-panel detectors (FPD), using which even a single “half rotation”, or more precisely a rotation through >180°+fan angle, enables informative CT views of limited volumes to be reconstructed by cone-beam computed tomography (CBCT). On the other hand, use is made of multi-row detectors, with multiple rotations around the patient and continuous advance, which can show the complete body of a patient tomographically by so-called spiral CT.
A serious problem is the scattered radiation produced in the exposed object, generally the patient, which increases with the area of the detectors and with the corresponding simultaneous increase in volume through which the radiation passes. Here, the intensity of the scattered radiation can reach the same order of magnitude as the unscattered direct primary radiation, and in extreme cases can indeed outweigh it. The consequence is a distortion in the quantitative reconstruction of the tissue thicknesses which, typically, in the case of a homogeneous cylindrical body for example, can be seen as a continuous darkening from the edge towards the center. This effect is called “cupping”. This cupping can have a very detrimental effect on the ability to recognize small pathological structures with low contrast, since errors of a deviation of several 100 HU can arise between the edge of the object and its center. 1 HU (=Hounsfield Unit) corresponds to 1/1000 of the density of water.
With conventional CTs with single row or multi-row detectors, the scattered radiation is effectively suppressed by collimation. This is done, for example, by limiting the thickness of the exposed layer in the case of single row detectors, or the volume of the layer for multi-row detectors, by axial collimation towards the axis of rotation. Use can also be made of encapsulation, or collimation plates arranged between the sides of neighboring detector elements, as appropriate, that is lateral collimation.
In the case of flat-panel detectors, such as are mainly used in modern C-arm devices, the radiation is, not least on grounds of radiation protection, restricted by collimators to the measurement area—in general the detector area, in many applications even to just a part of the area. However, the scattered radiation which is produced in the patient cannot be suppressed by these measures. Otherwise, there is no possibility of lateral collimation at the flat-panel detector.
It is also possible to use antiscatter grids mounted in front of the detector input surface. Though their benefits for CBCT imaging are the subject of controversial discussions, as emerges from the paper by J. H. Siewerdsen, D. H Moseley, B. Bakhtiar, S. Richard, D. A. Jaffray: “The influence of antiscatter grids on soft-tissue detectability in cone-beam computed tomography with flat-panel detectors”, Med. Phys. 31(12), December 2004, 3506 to 3520, their use is to be recommended at least when there is a large proportion of scattered radiation. In general however, the reduction in scattered radiation by antiscatter grids is insufficient, so that additional scattered radiation correction methods are necessary.
In the paper by R. Ning, X. Tang, D. L. Conover: “X-ray scatter suppression algorithm for cone beam volume CT”, Proc. SPIE, Vol. 4682, 2002, 774 to 781, it was suggested that the scattered radiation should be measured and corrected in a few projection directions. For the measurement of the scattered radiation this requires that equipment is available with which a “beam-stop” carrier plate can be moved into the path of the radiation between the X-ray source and the patient, close to the patient. However, the additional measurement procedure thus required seems to be unacceptable in the normal clinical workflow.
In the paper by M. Zellerhoff, B. Scholz, E.-P. Rührnschopf, T. Brunner: “Low contrast 3D reconstruction from C-arm data”, Proceedings of SPIE, Medical Imaging 2005, Vol. 5745, pages 646 to 655, a purely computational scattered radiation correction algorithm is suggested. This algorithm is indeed relatively fast, but the lengthening of the complete reconstruction time is not negligible. Thus, for 500 projection images with a size of 1024×1024 pixels, the algorithm still requires about 1 minute on a PC with a clock speed of 3 GHz.